High-precision drug delivery by dual-domain ocular device

ABSTRACT

The present invention relates to a nanocomposite ocular device that can release drugs within a close distance to the ocular surface and provide controlled and sustained release of the drug at a constant rate. The device can achieve both optical and medical functions. The device comprises a drug, one or more reservoir domains, and a barrier layer configured to block the drug diffusion paths from the reservoir domain to the ocular surface in the eye of the subject, wherein the drug partitions between the reservoir domain and the barrier layer, and the equilibrium drug solubility in the reservoir domain is at least five folds higher than that in the barrier layer.

This application is a continuation of U.S. application Ser. No.16/370,407, filed Mar. 29, 2019; which is a continuation ofPCT/US2017/054505, filed Sep. 29, 2017; which claims the benefit of U.S.Provisional Application No. 62/402,509, filed Sep. 30, 2016. Thecontents of the above-identified applications are incorporated herein byreference in their entirety.

FIELD OF THE INVENTION

The present invention relates to a drug-eluting nanoengineered oculardevice comprising a drug, one or more reservoir domains, and a barrierlayer configured to block the drug diffusion paths from the reservoirdomain to the ocular surface in the eye of the subject, wherein the drugpartitions between the reservoir domain and the barrier layer, and theequilibrium drug solubility in the reservoir domain is at least fivefolds higher than that in the barrier layer.

BACKGROUND OF THE INVENTION

The past decade has seen a great increase in patients wearingprescription contact lenses and color contact lenses in order to correctvisual sensation and change appearance of eyes. The advancement insilicone-hydrogel composite technology has improved both the wateraffinity and oxygen permeability of a contact lens and transformed itinto a breathable and hydrophilic soft tissue. Integrating thishydrophobic-hydrophilic nanocomposite technology with other advancementsin materials and medical nanotechnology, we can develop a new generationof nanoengineered ocular device that has the same optical performanceand ease of use as a contact lens, but with many more medicalapplications beyond simple vision correction. Because such an oculardevice is in constant contact with tear fluids at the most observablelocation of the eye, and is easily applied and removed, it can functionas the most convenient implantable medical device for diagnostics, drugdelivery, wound healing, and in particular, as a groundbreaking tool inocular disease therapy.

Delivering a drug to the eye at high therapeutic efficacy is challengingdue to the eye's complexity, including high sensitivity to foreignstimuli, protective mechanisms (e.g., blinking, tears), and severaloverlaying diffusion barricades (Short B, Soc Toxicol Pathol 2008; 36:49-62). The two main alternatives to implants have serious limitationsand drawbacks as follows.

1. Topical application of eye drops results in a sudden, burst dosagethat is difficult to control, and further hindered by tear drainagealong with corneal and scleral barriers to the point that only about 5%of functional ingredients deliver to the target, and

2. Intravitreal injection, while direct and efficacious, also gives aburst dosage that raises intraocular pressure, and requires repeatedadministration that bears risks to the patient's eye includinghemorrhage, cataract, retinal wounds and even retinal detachment.

There is a need for an ocular device that can effectively deliver a drugto the eye and release the drug at a dosage rate that is preciselycontrollable and sustainable.

BRIEF DESCRIPTION OF THE DRAWING

FIGS. 1A and 1B are schematics that detail the position and morphologyof drug reservoir domain(s) and barrier layer in a dual-domain contactlens in relevance to the post-lens-tear-film.

FIG. 2 is a schematic showing how drug partition between the reservoirand barrier layer maintained at the same drug activity (fugacity). Thedecrease in drug concentration in a long-term delivery, the tapering offof the drug activity at the barrier boundary due to the concentrationdrop in reservoir domain throughout the delivery period will cause verylittle concentration change in barrier layer because of the barriersaturation. Drug concentration in the barrier layer is always near aconstant despite the drug concentration variations within a range, X^(L)to X^(U), which leads to a constant-rate drug release throughout thedelivery.

FIG. 3 is a schematic and supporting mathematical equations showing howconstant drug delivery rate due to solubility saturation of barrierlayer by the drug in the reservoir.

FIG. 4 is a schematic illustrating how a reservoir matrix is surroundedby a barrier layer to regulate its drug release rate by solubilitysaturation at its boundary with the barrier. The porosity channels arecreated to facility drug loading by a soaking method.

FIG. 5 is a plot showing the cumulative release of bimatoprost from acontact lens under sink conditions following loading the contact lenswith bimatoprost hexanol solution.

FIG. 6 is a plot showing the cumulative release of bimatoprost from acontact lens under sink conditions following loading the contact lenswith bimatoprost dissolved in solutions with different hydrophilicity.

FIG. 7 is a plot showing the cumulative release of the drug cyclosporineA (CSA) from a contact lens under sink conditions after loading usingethanolic solutions of CSA.

FIG. 8A is a plot showing the cumulative release amount under sinkconditions of bimatoprost from silica-PMAA nanoparticles in phosphatebuffer with different pH values (pH 2.5 and pH 7.5). FIG. 8B is a plotshowing the cumulative release percentage of bimatoprost, expressed as apercentage of total from bimatoprost loaded silica-PMAA nanoparticles inphosphate buffer with different pH values (pH 2.5 and pH 7.5).

FIG. 9 is a plot showing the cumulative drug release of cyclosporine Aloaded silica-PMAA nanoparticles under sink conditions into phosphatebuffer with different pH values (pH 2.5 and pH 7.5).

FIG. 10 is a schematic that details the nanoengineered ocular device'sposition on the eye, and the function of the drug reservoir domain in awedge configuration.

DETAILED DESCRIPTION OF THE INVENTION Definitions

A “drug” as used herein, refers to a molecule having an activity tocause a physiological change in a subject, such as a pharmaceutical drugor a nutrient. For example, a drug can be a prescription drug asregulated by FDA, including, but not limited to, antibiotics(fluoroquinolone), glaucoma drugs (bimatoprost), dry eye drugs(cyclosporine), macular degeneration drugs (bevacizumab), or steroids(difluprednate); or an over-the-counter drug as regulated by FDA,including, but not limited to, artificial tears (carboxymethylcellulose)or contact lens solution; or a macronutrient such as a carbohydrate,protein, or fat (omega-3 fatty acids).

The “solubility parameter” of a substrate is an approximate measure ofmolecular interaction between identical molecules. When two differenttypes of molecules are mixed in one solution, the square of thedifference of the two solubility parameters reflects the energy penaltyof mixing. Thus, the greater the difference in solubility parameters,the lesser the solubility, which is the basis of the principle “likedissolves like.”

The “hydrophobic/hydrophilic” ratio is a measure of a material'shydrophobicity based on its solubility and partition between hydrophobicoctanol and hydrophilic water. It is normally referred to a materialconstant log P of octanol/water, where P is a partition coefficient of asoluble component between octanol and water (Sangster J, J Phys Chem RefData 1989; 18:No. 3). A positive log P value means hydrophobic and anegative value means hydrophilic. A differential of log P less than 0.5between two components means they are of a similar (alike) nature, whilea differential of log P greater than 1 means the two components are ofopposite (dislike) nature. Thus. the difference in log P has the sameimplication as the difference in solubility parameters. However, bothmeasures are just qualitative, not quantitative, descriptions of thesolubility behavior of the mixing components.

The inventors of the present invention have discovered a nanocompositeocular device that can release one or more drugs within a close distance(about 10μ) to the cornea and at a dosage rate that is preciselycontrollable and sustainable. This device has similar optical quality(i.e., transparent) to a vision-corrective contact lens. The device canbe non-surgically (i.e. by hand, or with use of a tool) overlaid eitherin contact with the tear film above the cornea or sclera, or directlyonto the cornea or sclera. The device can achieve both optical andmedical functions. An optically and medically combined function meansthe portion of the device covering the cornea must be highly transparent(transmission rate over 90%), with or without vision corrective power,while the device achieves one or more desired medical function (i.e.,drug delivery, tissue engineering, sensing, diagnosing, orthokeratology,environmental responses, etc.).

The present invention is directed to a transparent ocular device fordelivering one or more drug to the eye of a subject. The devicecomprises: (i) at least a drug, (ii) one or more reservoir domains, and(iii) a barrier domain of layer configuration (or a barrier layer) toblock the drug diffusion paths from the reservoir domain to the ocularsurface in the eye of the subject; wherein the drug partitions betweenthe reservoir domain and the barrier layer, and the equilibrium drugsolubility in the reservoir domain is at least five folds higher thanthat in the barrier layer, such that when the device is placed directlyonto the ocular surface or in contact with the tear film above theocular surface in the eye, the device provides controlled and sustainedrelease of the drug at a constant rate. In one embodiment, the device isa drug-delivery optical lens optionally having refractive power tocorrect myopia, hyperopia, astigmatism, presbyopia and/or other opticalwave guiding power. For example, the device is a contact lens, scleralens, or ortho-K lens. In another embodiment, the device is atissue-engineering scaffold, wound-dressing gel, or an ocular bandage,etc. The device does not need to attach to another ocular medical device(such as an optical lens or a contact lens) to achieve a medicalfunction.

The present invention incorporates the following features into thedesign of an ocular drug delivery device so that it is capable ofdelivering a drug at a near constant rate (zero-order kinetics), with ahigh-precision dosage and extended duration (a week or longer).

-   -   (a) The drug is primarily dissolved into a reservoir domain        which is composed of components with a solubility parameter        similar to the drug in order to achieve a high drug solubility.    -   (b) The drug-embedded reservoir domain is either encapsulated as        isolated domains within a barrier layer (FIG. 1A), or it is        formed as a separate layer in a series contact with a barrier        layer (FIG. 1B). The components of the barrier layer are        incompatible to the drug so that the drug has an inherently low        solubility within the barrier—leading to a drug's partition        predominantly in the reservoir and always near a saturation in        the barrier. The in-series arrangement of drug's diffusion path        from a reservoir, through a barrier of extremely low drug        solubility, and into the post-lens-tear-film assures the drug        release rate from reservoirs be completely controlled by the        drug's saturation solubility in the barrier.    -   (c) The drug solubility partition coefficient between the        reservoir domain and the barrier layer is at least 5 to 1,        preferable 10 to 1, 30 to 1, 50 to 1, or 100 to 1. Due to the        solubility difference between the reservoir domain and the        barrier layer and an excessive drug population within the        reservoir domain, the device can consistently discharge at a        near-constant rate with high-precision dosages dictated by the        saturation solubility in the barrier layer for a long period of        time.        Solubility Saturation in the Barrier Layer

The substantial difference of drug solubility in the reservoir domainand barrier layer is illustrated in FIG. 2. The solubility behaviorgenerally follows the rule of like-dissolve-like. A hydrophobic drug hasa higher solubility in a hydrophobic medium but very limited solubilityin a hydrophilic environment. For a definition of the hydrophobic andhydrophilic nature of a component, one can use the log P octanol/wateras a reference point, where P is a partition coefficient of a solutebetween octanol and water. A positive log P value means hydrophobic anda negative log P value means hydrophilic. A differential of log P lessthan 0.5 between two components means they are of a similar (alike)nature, while a differential of log P greater than 1 means the twocomponents are of opposite (dislike) nature.

The ocular device of the present invention is a dual domain device,which contains a reservoir domain with a similar (alike) log P to thedrug and a barrier layer with a very different (dislike) log P to thedrug so that the drug has a substantially higher (e.g., ten times ormore) solubility in reservoir than in barrier. The reservoir domain isreferred to as the “drug-like domain” and the barrier layer as the“drug-dislike domain.” The large disparity in drug solubility of the twodomains has created a drug concentration window within which the changein drug's fugacity/activity (˜Exp(Δμ/kT), where Δμ=μ−μ° is the drug'schemical potential relative to a reference standard state) correspondsto a proportionally substantial concentration change in the reservoirdomain but almost constant concentration in the barrier layer due to thesaturation effect. When an ocular device is operating within this drugconcentration range in the reservoirs, the drug concentration in thebarrier layer remains almost constant (saturation effect) throughout theentire delivery period. For each specific drug delivery, the initialhighest drug concentration in reservoir, X^(U), is loaded according toclinically established efficacy and safety levels. The drug delivery isessentially at a constant rate due to the barrier saturation (detailselaborated in the next section) until the reservoir concentration dropsto a lower bound X^(L), below which the reservoir concentration is notsufficient to maintain barrier saturation anymore. Beyond this point,the drug delivery rate is no longer be a constant and, in fact,diminishes proportionally following the depletion of the residualreservoir drug level below X^(L). The lower a drug solubility in abarrier layer, the smaller the X^(L) of reservoir concentration level.For any specific drug, the ratio R_(x)=X^(U)/X^(L), defined as thedrug's reservoir/barrier saturation ratio, is higher with increasingdrug solubility difference between reservoir and barrier, while theratio 1/R_(x), indicates the last delivery portion not achieving theprecision of a constant rate.

Delivery Rate Controlled by the Barrier Layer

The drug's steady-state permeation rate thorough a barrier layer can beprecisely prescribed by the multiplication product of a diffusionconstant (a material property of the barrier) to the drug'sconcentration gradient in the barrier, the latter of which can beexpressed as

$\frac{C_{int} - C_{out}}{t},$where t is the thickness of a barrier. The barrier must be positioned toblock the drug diffusion paths from the reservoir domain to the tearfilm between the device and the ocular surface (e.g., see FIGS. 1A, 1B,and 2) so that it can throttle the drug release as the device'rate-determining step (i.e., the slowest rate in a series of diffusionpaths). In a steady state delivery, the drug permeation rate through thebarrier must equal the combined rates of corneal adsorption(J_(adsorption)) and tear turnover loss (J_(turnover)). Once reaching asteady-state, the permeation rate through the barrier remains nearly ata constant due the saturation effect at the boundary of reservoirdomains and the barrier layer (i.e. C_(in) is aconstant≈C_(saturation)). (FIG. 3)High-Precision Drug Delivery Rate and Dosage to Optimize CornealAdsorption

The dual-domain ocular device of the present invention separates thefunctions of drug storage and rate-determination into the respectivereservoir and barrier layers, and thus, can achieve the highestprecision in the control of rate and dosage of ocular drug deliveries.Eye drops in ocular treatment are known for low delivery efficacy asless than 5-10% are available to the cornea. Eye drop treatments ofglaucoma are almost completely empirical with only limited control overefficacy through dosing frequency and concentration. On the other hand,the nanoengineered ocular device of the present invention delivers drugsat a long-term, steady (near constant) rate and with the highestpossible precision. For each specific drug, we can combine in-vitrokinetics studies and animal models to first quantify the most effectivedosage (i.e. a steady-state concentration in the tear film between thedevice and the ocular surface matching an optimized corneal adsorptionrate), followed by fine-tuning the material properties of the dualdomains (e.g., hydrophobic/hydrophilic ratio, degree of crosslinking,barrier thickness, porosity etc.) to deliver the highest-efficacyconcentration throughout the treatment period using the device. Thisapproach substantially increases the drug's bioavailability and reducespotentially harmful systemic absorption (estimated to be at least 90%less than by topical eye drops), among many other deliveryinefficiencies.

Dual-Domain Design Maximizes Efficacy of an Ocular Device Drug Delivery

In one embodiment, the present invention utilizes the dual-domaintechnology and transforms the otherwise generic contact lens drugdelivery to a custom-designed, ocular device to deliver a specific drugwith optimal precision, efficacy, and duration. The dual-domain lensdelivery technology of the present invention can precisely andconsistently correlate the steady-state drug delivery rate to thematerial properties of a composite ocular device in terms ofhydrophilic/hydrophobic composition ratio of the device and thediffusion speed in the rate-determining barrier layer. Based on a drug'ssolubility and partition manners with different solvents and among lens'different domains, efficient drug loading processes can be achieved bysoaking the dual-domain ocular device of the present invention in amixture of solvents. Thus, the dual-domain ocular device, besidesoffering higher quantitative precision than generic lens delivery, canbe fabricated with custom-designed composite composition to make itsdelivery exactly matching the most preferred corneal adsorption rate(provided by animal model in each specific drug case), while make thedevice's drug loading process sufficiently precise and batch-consistentfor mass productions and clinical tests. The high-precision drug loadingas well as delivery from a specially designed dual-domain lens are keyfactors for the commercialization of ocular delivery device.

The permeation rate of a drug through a substrate barrier is generallydetermined by the product of the drug's solubility (S) and diffusivity(D) in the substrate; i.e. Permeability=S·D. A drug's solubility in asubstrate is normally increased with a higher loading concentration atthe substrate boundary (normally expressed in terms of activity which isreferenced to a standard state and equals to the concentrationmultiplied by an activity coefficient) until it levels off to a constantnear a saturation solubility. Beyond this saturation point, any furtherincrease of concentration at the substrate boundary will not change theconcentration within the substrate and thus, will not affect the drugpermeation rate in substrate given that the drug diffusion coefficientis a constant. Covering a drug reservoir externally with a continuousbarrier layer that has a low saturation solubility and constantdiffusion coefficient (a layer with no swelling and no porosity) canachieve a zero-order release because the drug permeability in thebarrier is always a constant at the saturation point.

The practice of a dual-domain design of the present invention isillustrated in FIG. 4. A drug is loaded within a reservoir polymerdomain that has a low matrix diffusion coefficient, D_(m), but is highlycompatible to the drug (for example, a hydrophobic drug dissolved withina hydrophobic polymer matrix) leading to a high S_(m). The polymermatrix contains open-channel pores from which the drug can be quicklyequilibrated among nearby reservoir domains. A barrier coatingcontaining no porosity and with very low drug solubility is applied tocover the drug-loaded nanopore polymer matrix of which the drugconcentration in the reservoirs, or pores are much higher than thesaturation solubility in the barrier, S_(b), so that the resultantleaching rate is completely regulated by the barrier permeability,P_(b)=S_(b)·D_(b). For example, when the reservoir and barriercomponents are chosen to have a saturation ratio of 10 (i.e.R_(x)=X^(U)/X^(L)˜10), an estimated 90% of the drug,

${\frac{X^{U} - X^{L}}{X^{U}} \times 100\%},$is released in zero-order rate and the last 10% (100%/R_(x)) in adiminishing rate proportional to the remaining drug level.

In a dual-domain device of the present invention, the solubilitydifference defined by a drug's reservoir/barrier saturation ratio ofR_(x) is at least 5 to 10, preferably with even higher multiple up to100. The barrier layer in general requires a diffusive thickness of0.5-100 μm, preferably 10-100 μm, or 30-100 μm, which inverselyproportion to the layer's diffusion constant. The reservoir domain andthe barrier layer preferably, but not necessarily, are in direct contactwith each other, or connected with pore channels. The connection throughopen channels facilitates drug adsorption by the reservoir domain in asolution drug loading process.

In one embodiment, the drug is hydrophobic, the reservoir domain is madeof one or more hydrophobic components, and the barrier layer is made ofone or more hydrophilic components.

In another embodiment, the drug is hydrophilic, the reservoir domain ismade of one or more hydrophilic components and the barrier layer is madeof one or more hydrophobic components.

The general rule of choosing the reservoir and barrier materialcomponents for a specific drug is to make the log P of the reservoirdomain very close to log P of the drug. Exactly how close of the log Pof reservoir to the drug depends on how much drug solubility is requiredin the reservoir domain to achieve the desired clinical efficacy. Thedifference in log P between reservoir and barrier layer is preferredhigher than 1, more preferably 2 to 3, or even higher. Using a mixturein a domain is acceptable and sometimes necessary to maintain componentscompatibility and transparency. There are methods to obtain log P of acomposite. In order to also maintain other desired properties of atransparent lens, such as oxygen permeability, water affinity, andflexibility, the hydrophobic and hydrophilic components may be chosenbased on ingredients already successfully utilized by hydrogel,silicone-hydrogel, and rigid-gas permeable lenses as listed below inTables 1A-1C.

For example, the hydrophilic components of the reservoir domain/barrierlayer can be selected from any of the hydrophilic components listed inTable 1A, 1B, or 1C, and the hydrophobic components of the reservoirdomain/barrier layer can be selected from any of the hydrophobiccomponents listed in Table 1A, 1B, or 1C,

TABLE 1A Common materials used to make hydrogel soft contact lensHydrophilic component Hydrophobic component 2-Hydroxyethyl methacrylate(“HEMA”) Methyl methacrylate (“MMA”) N,N-dimethylacrylamide (“DMA”)Isobutyl methacrylate N-vinyl-2-pyrrolidone (“NVP”) Pentyl methacrylate4,4-Dimethyl-2-vinyl-2-oxazolin-5-one Cyclohexyl methacrylateMethacrylic acid (“MAA”) Lauryl methacrylate N-(Hydroxymethyl)acrylamideN-[3-(Dimethylamino)propyl]methacrylamide Ethylene glycol dimethacrylate

TABLE 1B Common materials used to make silicone-hydrogel soft contactlens Hydrophilic component Hydrophobic component 2-Hydroxyethylmethacrylate 3-[Tris(trimethylsiloxy)silyl]propyl methacrylateN,N-dimethylacrylamide (“TRIS”) N-vinyl-2-pyrrolidone3-Methacryloxy-2-hydroxypropoxy 4,4-Dimethyl-2-vinyl-2-oxazolin-5-(propylbis(trimethylsilyloxy)methylsilane one (“SIGMA”) Methacrylic acidFluorosiloxane macromer 2-(Methacryloyloxyethyl)-2-Mono-(3-methacryloxy-2-hydroxypropyloxy)propyl (trimethylammonioethyl)phosphate terminated, mono-butyl terminated Ethylene glycoldimethacrylate polydimethylsiloxane Poly(N-vinyl pyrrolidone) (“PVP”)Mono-methacryloxypropyl terminated Triethyleneglycol dimethacrylatepolydimethylsiloxane

TABLE 1C Common materials used to make rigid gas permeable (“RGP”)contact lens Hydrophilic component Hydrophobic component 2-Hydroxyethylmethacrylate Methyl methacrylate Methacrylic acid Trifluoroethymethacrylate Ethylene glycol dimethacrylate Hexafluoroisopropylmethacrylate Neopentyl glycol dimethacrylate3-[Tris(trimethylsiloxy)silyl]propyl methacrylatePentacontamethyl-α,ω-bis-(4-methacryloxybutyl) pentacosasiloxane1,3-Bis(3-methacryloxypropyl) tetrakis(trimethylsiloxy)disiloxane Bishexafluroisopropyl itaconate Fluoro-siloxanyl styrene

All present contact lenses are made from acrylic oligomers, orprepolymers by a thermal curing (100° C.-120° C.) process followed by anextensive cycle of washing and sterilization. A drug can be incorporatedin a lens either before, or after the lens fabrication process. Thepre-fabrication loading of a drug by mixing with lens ingredients iseasy, but protecting a drug from thermal degradation, or loss during thewashing cycle could be challenging. By examining a drug's solubility andpartition coefficient among the relevant solvent(s) and lens components,we can design a post-fabrication drug loading scheme to precisely loadthe desired drug dosage in a finished lens. One key is to utilize theprefabricated porosity channels (20-40% porosity) to accelerate theloading kinetics. The composition and conditions of the soaking solutioncan also be manipulated (for example, utilizing temperature, orsupersaturation) to promote the drug's faster adsorption into thereservoirs.

In one embodiment, the invention is directed to a drug-eluting contactlens, which provides controlled and sustained release of the therapeuticagent when contacting the eye, by incorporating

-   -   (1) a hydrophobic drug (for example, bimatoprost, log P˜3.2) in        reservoirs made of one or more hydrophobic components listed in        Table 1B, embedded within a continuous barrier made of one or        more hydrophilic components listed in Table 1B, along with        chemical additives for curing and crosslinking and 20% inert        solvents for porosity creation;    -   (2) or, a hydrophilic water-soluble drug (for example timolol        maleate, log P˜1.4) in reservoirs made of one or more        hydrophilic components listed in Table 1B and cover the        reservoirs with a continuous barrier composed of one or more        hydrophilic components listed in Table 1B; along with chemical        additives for curing and crosslinking and 20% inert solvents for        porosity creation;        -   such that the reservoir/barrier saturation ratio is ten or            higher and the barrier thickness (10-100 microns) adjusted            to control the constant release rate to the desired level            (for example, 20 μg/day for a week).

In another embodiment, a drug-eluting contact lens can be fabricated bypre-forming, or pre-synthesizing drug-philic reservoir components indomains of block-copolymers, nanoparticles, liposomes, micelles, etc.,and in shapes of needles, rods, disks etc., followed by incorporatingthese reservoir domains within a continuum of drug-philic barriercomponents cured by free radical polymerization, condensation reaction,gelation, or drying. The drug can be loaded by solution soaking orsolvent exchange in making reservoir particles, or solution soaking ofthe finished devices.

In one embodiment, a drug-eluting contact lens can be fabricated byreacting the drug-philic reservoir components to a form of a condensedstate of a polymer, a gel, a block-copolymer, a nanoparticle, asurface-modified nanoparticle, a connected nanodomain, and aninterpenetrating network. The drug in the reservoir layer may beentrapped within a porous nanoparticle to prevent unintentional burstrelease of the drug from the reservoir layer. For example, a hydrophobicdrug can be dispersed into the pores of a hydrophilic nanoparticle madeof alginate and/or chitosan, and the interpenetrating network ofalginate or chitosan can function as a hydrophilic barrier either forslowing down a drug release, or preventing its burst release as an extrasafety control measure.

Using several additional options, we address the safety control of thedrug delivery device. Because the delivery rate is completely regulatedby the low permeation rate within the barrier layer, both the geometryof and the permeation rate within the reservoir layer can be flexiblymodified without compromising the delivery kinetics so long as thereservoir layer is in series with a barrier layer and its intrinsic drugpermeation within the reservoir does not fall short of the drugpermeation rate in the bather. The geometry of the drug “reservoir” zonecan be changed from a flat film to other morphologies such asnanoparticles for facilitating drug incorporation, domain inclusion,device fabrication, or the drug delivery function as long as its rate isstill regulated by a barrier in series.

The drug-eluting ocular device of the present invention optionallyincludes multiple safety measures to prevent burst failure. Since themajority drug content resides within the reservoir layer, theflexibility of changing morphology and geometry of the reservoir domaincan add several safety measures to prevent an accidental storage burstfailure. The inventors have developed three additional safety options toeliminate any catastrophic burst failure: (a) hardening the reservoirdomain to lower drug diffusion within; (b) entrapping the drugbimatoprost into silica-alginate nanoparticles as an extra barrier; and(c) making the drug reservoir into a wedge-shaped ring with the narrowerend facing cornea and functioning as a safety valve. By doing so, thedevice has another safety valve (narrowing the channel) against leakinginto the tear film between the device and the ocular surface. In oneembodiment, the reservoir domain is configured into a wedge-shapedgeometry with a narrow end facing the ocular surface when placed in theeye and a broad end facing away from the ocular surface (FIG. 10) whenplaced in the eye, to prevent any accidental over dosing in the tearfilm between device and eye. In another embodiment, the reservoir domainis configured into a wedge-shaped geometry with the narrow end pointingat a specific area of the ocular surface when placed in the eye toachieve a targeted, location-specific delivery of the drug. Further, thedevice ensures that even if all other safety measures fail, the leakagewill be predominantly through the top surface of the device where tearlacrimation would wash away the majority of overflow, similar to atopical eye drop application. Such a novel structural design alsoenables the pinpoint precision of a location-specific drug deliveryarea.

The drug-eluting ocular device of the present invention targets drugdelivery to a specific location. We utilize the location of the drugloading zone and the advantage of a short, but direct delivery route toachieve a localized delivery targeted at a specific spot on the surfaceof the cornea. Such a location-specific delivery is possible only by adirect delivery to ocular surface because the device is centered at alocation relatively stationary with respect to the cornea, exceptoccasional lateral movements. A glaucoma, or other, drug can bedelivered to a specific location near the edge of the cornea to targetprostaglandin receptors in the trabecular meshwork, sclera, and ciliarymuscle. Because of the short distance of diffusion (˜20 μm) relative tothe dimension of the target (a few millimeters), the drug permeation ininterior tear film is considered to be unidirectional toward the target(FIG. 3). The targeting of a specific location around the cornea fordrug delivery can benefit ocular treatments other than glaucoma as well.For example, the corneoscleral rim of the eye (about 12 mm in diameterof the center) represents a region with unique anatomical properties dueto the location of adult corneal epithelial stem cells in the corneallimbus. These stem cells are essential for the healing and regenerationof the corneal epithelium and for preventing scar tissue growth and atargeted delivery of growth factors and/or nutrients will be beneficialto corneal repair or tissue regeneration-based treatment.

The following examples further illustrate the present invention. Theseexamples are intended merely to be illustrative of the present inventionand are not to be construed as being limiting.

EXAMPLES

The following examples show dual-domain contact lens of the presentinvention having constant-rate drug delivery. All drug release examplesare carried out under sink conditions (using 0.5 ml simulated tear fluidfor each lens and replaced by 0.5 ml at each time point) to simulate thetear exchange mechanism in the real case. All examples have shown aquicker release for a few hours followed by a constant-rate release forover many days. The initial rapid drug release is attributed to theleaching of drug loading solution remaining in the pore channels. Therapid release stage can be eliminated by a post-loading solvent exchangeprocess. This loading process and the subsequent release process allutilize the drug solubility equilibrium among all components that notonly provides the most precise deliverable dosage, but also can preventpremature drug release in storage by creating a drug solubilitysaturation in the storage medium for long-term storage.

Example 1

This example provides an example of fabrication of silicone-hydrogeldual-domain contact lens.

The soft contact lenses were prepared by polymerizing 2-hydroxyethylmethacrylate (HEMA 8%) along with highly oxygen permeable siliconemonomer SIGMA (30%), N.N-dimethylacrylamide (DMA ˜30%), n-hexanol (23%to create porosity), free-radical thermal initiator2,2′-azobis-(2-methylpropionitrile) AIBN, polyvinylpyrrolidone K-90 (5%for wetting) and ethylene glycol dimethacrylate (EGDMA, 1% forcrosslinking). The mixture of monomer mixed was injected into apolypropylene contact lens mold. After curing at 120° C. for 2 hours,contact lenses were washed with isopropanol and water mixture (50%, v/v)to remove initiator residue and unreacted monomers. For nanoparticleloaded contact lenses, nanoparticles were added into pre-cured monomersolution subsequent to polymerization.

Example 2

Silicone-hydrogel lens loaded with bimatoprost (a hydrophobic drug, logP˜3.2) by soaking lens in hydrophobic solvent hexanol (log P=2.03): Forthe drug loading process, three silicone hydrogel soft contact lenseswere soaked in 0.5 ml bimatoprost hexanol solution with a concentrationof 50 mg/ml for two days; another three soft contact lenses were soakedin 0.5 ml bimatoprost hexanol solution with a concentration of 25 mg/mlfor two days. Each group of lenses was rinsed in 5 ml water for 30seconds.

For the in vitro release condition, each lens was placed in 0.5 mlphosphate buffer (pH 7.5, 10 mM), at predetermined time intervals, 0.5ml release medium was withdrawn and replaced with fresh phosphate bufferto simulate the tear exchange sink condition. The amount bimatoprostreleased into the phosphate buffer medium was measured using a UV/VISspectrophotometer at a wavelength of 210 nm. Concentrations and massesof released bimatoprost at each kinetic time point were calculated basedon a calibration curve prepared with known bimatoprost concentrations(R²>0.99).

The cumulative mass release profile of the release of bimatoprost overtime from the contact lens is plotted in FIG. 5 and shows a constantrate of sustained release of the drug for 12 days. Hexanol is ahydrophobic solvent, log P˜2, it can dissolve a high amount ofbimatoprost and thus, carry a high drug dosage into the siliconehydrophobic reservoir domain of the contact lens or ocular device.Hexanol loading at high bimatoprost concentration can achieve constantrate release due to saturation of the hydrophilic barrier domain formedfrom the polymerization of HEMA and DMA components of the contact lens.

Example 3

Silicone-hydrogel contact lens loaded with bimatoprost by soaking lensin a medium with different hydrophilicity—Water and Water/Isopropanol(IPA) mixture (50% v/v): For the drug loading process, three softcontact lenses were soaked in 1 ml bimatoprost water/isopropanol (50%v/v) solution with a concentration of 0.5 mg/ml for two days; anotherthree soft contact lenses were soaked in 1 ml bimatoprost water solutionwith bimatoprost concentration of 0.5 mg/ml for two days. Each group oflenses was rinsed in 5 ml water for 30 seconds.

For the in vitro release condition, each lens was placed in 0.5 mlphosphate buffer (pH 7.5, 10 mM), at predetermined time intervals, 0.5ml release medium was withdrawn and replaced with fresh phosphate bufferto simulate tear exchange sink condition. The amount bimatoprostreleased into the phosphate buffer medium was measured using a UV/VISspectrophotometer at a wavelength of 210 nm.

The cumulative mass release profile of the release of bimatoprost fromthe contact lens is plotted in FIG. 6. The contact lens in water loadingcondition showed slightly higher release amount compared to the siliconehydrogel contact lens in isopropanol (50% v/v) loading condition.Isopropanol is less hydrophobic (Log P=0.05) then hexanol, but morehydrophobic than water. The isopropanol and water mixture can dissolvemore bimatoprost than pure water. At the same drug loading (0.5 mg/ml)concentration, the drug's activity (fugacity) in water/isopropanolmixture is higher than in water and can load a higher amount ofbimatoprost into the hydrophobic reservoir.

Example 4

Silicone-hydrogel lens loaded with cyclosporine A (a hydrophobic drug)by soaking in ethanol solution: For the drug loading process, three softcontact lenses were soaked in 2 ml cyclosporine A ethanol solution witha concentration of 3 mg/ml for two days. Each group of lenses is rinsedin 5 ml water for 30 seconds.

For the in vitro release condition, each two lenses were placed in 0.5ml phosphate buffer (pH 7.5, 10 mM), at predetermined time intervals,0.5 ml release medium was withdrawn and replaced with fresh phosphatebuffer. The amount cyclosporine A released into the phosphate buffermedium was measured using a UV/VIS spectrophotometer at a wavelength of205 nm. Concentrations and masses of released cyclosporine A at eachkinetic time point were calculated based on a calibration curve preparedwith known cyclosporine A concentrations (R²>0.99). The cumulative massrelease profile of the release of cyclosporine A is plotted as afunction of time in FIG. 7. We used ethanol (Log P=0.05) to improve thesolubility in loading solution and obtain a constant rate release up toat least 40 days. Cyclosporine A is a hydrophobic (Log P=4.3) nonchargedcyclic peptide that confers a high rigidity to the structure and thedrug has very low aqueous solubility of 23 μg/mL at 25° C. (Miyake K, JPharm Sci 1999; 88:39-45). This hydrophobic drug has high affinity forthe silicone hydrophobic reservoir domain of the contact lens while thehydrophilic domain formed from the polymerization of HEMA and DMAcomponents is an effective barrier to control the release of the drug.

Example 5

Silica-poly(methacrylic acid) (PMAA) nanoparticles were prepared using awater-in-oil microemulsion system: Initially, 20 mL n-hexanol wasdissolved in 60 mL cyclohexane, followed by addition of 6 mL aqueousPMAA (3% w/w) solution. After 5 minutes, 25 mL Triton™ X-100 non-ionicsurfactant was added dropwise until the stirred solution becameoptically transparent. After 10 minutes of vigorous stirring, 500 μLaqueous ammonia solution (29% w/w) was added dropwise, followed byaddition of 1.5 mL tetramethyl orthosilicate. The reaction mixture wasstirred for 24 hours at 23° C. before addition of 100 mL acetone tobreak the stability of the microemulsion. The nanoparticles wererecovered by centrifugation (4500 rpm, 10 minutes) and then were washedthree times with isopropanol and deionized water to remove the excesssurfactant and cosurfactant.

Bimatoprost loaded silica-PMAA nanoparticles: Bimatoprost was firstdissolved in ethanol to make a 5 mg/mL solution. Then silica-PMAAnanoparticles were added into 100 μL of this solution. The mixture wasincubated at 23° C. for 5 hours and dried in an oven at 80° C. for 12hours. The drug loaded silica-PMAA nanoparticles were washed twice withphosphate buffer solution (10 mM, pH 2.5) to remove surface adsorbedbimatoprost.

In vitro release study: Bimatoprost loaded silica-PMAA nanoparticleswere suspended in 0.5 mL phosphate buffer (10 mM, pH 2.5) and phosphatebuffer (10 mM, pH 7.5) at 23° C. At predetermined time intervals, allrelease samples were centrifuged (14800 rpm) for 10 minutes, and 0.5 mLsupernatant was withdrawn and replaced with 0.5 mL fresh release medium.The amount bimatoprost released into the phosphate buffer medium wasmeasured using a UV/VIS spectrophotometer at a wavelength of 210 nm.Concentrations and masses of released bimatoprost at each kinetic timepoint in all release media were calculated based on a calibration curveprepared with known bimatoprost concentrations (R²>0.99). After therelease study, the silica-PMAA nanoparticles were dissolved in 1 mL NaOH(1 M) to quantify the bimatoprost left inside the nanoparticles. Thepercentage of cumulative released bimatoprost was then calculated.

The cumulative mass release profile of the bimatoprost released overtime is plotted in FIG. 8A and the corresponding percentage of thebimatoprost released over time is plotted in FIG. 8B. In this example,silica-PMAA nanoparticles showed pH-dependent release ability for loadedbimatoprost, more bimatoprost was released with the higher pH condition.This demonstrates feasibility of constant-rate release in response tochanges in environmental condition (pH). Bimatoprost was loaded directlyinto the nanoparticles and adheres on the pore interfaces after drying.The drug release was directly from the nanoparticles where the simulatedtear fluid in the pore channels served as the effective barrier layerbecause of saturation of the hydrophobic drug.

Example 6

Silica-PMAA nanoparticles were prepared using a water-in-oilmicroemulsion system: Initially, 20 mL n-hexanol was dissolved in 60 mLcyclohexane, followed by addition of 6 mL aqueous PMAA (3% w/v)solution. After 5 minutes, 25 mL TRITON® X-100 non-ionic surfactant wasadded dropwise until the stirred solution became optically transparent.After 10 minutes of vigorous stirring, 500 μL aqueous ammonia solution(29% w/w) was added dropwise, followed by addition of 1.5 mL tetramethylorthosilicate. The reaction mixture was stirred for 24 hours at 23° C.before addition of 100 mL acetone to break the stability of themicroemulsion. The nanoparticles were recovered by centrifugation (4500rpm, 10 minutes) and then were washed three times with isopropanol anddeionized water to remove the excess surfactant and cosurfactant.

Cyclosporine A loaded silica-PMAA nanoparticles: Cyclosporine A wasfirst dissolved in ethanol to make a 10 mg/mL solution. Silica-PMAAnanoparticles were added into 100 μL of this solution followed bysonication to obtain homogeneously dispersed nanoparticles. The mixturewas incubated at 23° C. for 5 hours and dried in an oven at 80° C. for12 hours. The drug loaded silica-PMAA nanoparticles were washed twicewith phosphate buffer solution (10 mM, pH 2.5) to remove surfaceadsorbed cyclosporine A.

In vitro release study: Cyclosporine A loaded silica-PMAA nanoparticleswere suspended in 1 mL phosphate buffer (10 mM, pH 2.5) and phosphatebuffer (10 mM, pH 7.5) at 23° C. At predetermined time intervals, allrelease samples were centrifuged (14800 rpm) for 10 minutes, and 0.5 mLsupernatant was withdrawn and replaced with 0.5 mL fresh release medium.The concentration of cyclosporine A in all release media was measuredusing a UV-VIS spectrophotometer at a wavelength of 205 nm.

The cumulative mass release profile of the cyclosporine over time isplotted in FIG. 9. Cyclosporine A is even more hydrophobic thanbimatoprost and consequently has even lower solubility in simulated tearfluid. The results showed constant rate release for much longer period(for at least 55 days) under both pH conditions due to the effectivebarrier effect of the simulated tear fluid in the pore channels of thesilica-PMAA nanoparticles.

The invention, and the manner and process of making and using it, arenow described in such full, clear, concise and exact terms as to enableany person skilled in the art to which it pertains, to make and use thesame. It is to be understood that the foregoing describes preferredembodiments of the present invention and that modifications may be madetherein without departing from the scope of the present invention as setforth in the claims. To particularly point out and distinctly claim thesubject matter regarded as invention, the following claims conclude thespecification.

What is claimed is:
 1. A transparent ocular device for delivering a drugto the eye of a subject, comprising: a hydrophobic drug; one or morehydrophobic reservoir domains made of one or more hydrophobic componentsselected from the group consisting of:3-[tris(trimethylsiloxy)silyl]propyl methacrylate (TRIS),3-methacryloxy-2-hydroxypropoxy (propylbis(trimethylsilyloxy)methylsilane (SIGMA), fluorosiloxane macromer,mono-(3-methacryloxy-2-hydroxypropyloxy)propyl terminated, mono-butylterminated polydimethylsiloxane, mono-methacryloxypropyl terminatedpolydimethylsiloxane, and poly(methacrylic acid) (PMMA); a barrier layermade of one or more hydrophilic components selected from the groupconsisting of: 2-hydroxyethyl methacrylate (HEMA),N,N-dimethylacrylamide (DMA), N-vinyl-2-pyrrolidone,4,4-dimethyl-2-vinyl-2-oxazolin-5-one, methacrylic acid,N-(hydroxymethyl)acrylamide, N-[3-(dimethylamino)propyl]methacrylamide,ethylene glycol dimethacrylate, polyvinylpyrrolidone (PVP), alginate,and chitosan, configured to block the drug diffusion paths from thereservoir domain to the ocular surface in the eye of the subject;wherein the drug partitions between the reservoir domain and the barrierlayer, and the equilibrium drug solubility in the reservoir domain is atleast five folds higher than that in the barrier layer, and when thedevice is placed directly onto the ocular surface or in contact with atear film above the ocular surface in the eye, the device providescontrolled and sustained release of the drug at a constant rate.
 2. Thetransparent ocular device of claim 1, wherein the one or morehydrophobic components are selected from the group consisting of: TRIS,SIGMA, fluorosiloxane macromer, and PMMA.
 3. The transparent oculardevice of claim 1, wherein the one or more hydrophilic components areselected from the group consisting of: HEMA, DMA, ethylene glycoldimethacrylate, PVP, alginate, and chitosan.
 4. The transparent oculardevice of claim 1, wherein the hydrophobic component is SIGMA, and theone or more hydrophilic components are DMA, ethylene glycoldimethacrylate, and/or PVP.
 5. The transparent ocular device accordingto claim 1, wherein the one or more reservoir domains are in a form of alayer.
 6. The transparent ocular device according to claim 1, whereinthe one or more reservoir domains are embedded within the barrier layer.7. The transparent ocular device according to claim 1, which is acontact lens, a scleral lens, an orthokeratology lens, or a cornealbandage.
 8. The transparent ocular device of claim 1, wherein thehydrophobic drug is bimatoprost.
 9. A method for treating glaucoma,comprising the step of applying the transparent ocular device of claim 8to the surface of the cornea of a subject in need thereof.
 10. Thetransparent ocular device of claim 1, wherein the hydrophobic drug iscyclosporine.
 11. A method for treating dry eye disease, comprising thestep of applying the transparent ocular device of claim 10 to thesurface of the cornea of a subject in need thereof.
 12. A transparentocular device for delivering a drug to the eye of a subject, comprising:a hydrophilic drug, one or more hydrophilic reservoir domains made ofone or more hydrophilic components selected from the group consistingof: HEMA, DMA, N-vinyl-2-pyrrolidone,4,4-dimethyl-2-vinyl-2-oxazolin-5-one, methacrylic acid,N-(hydroxymethyl)acrylamide, N-[3-(dimethylamino)propyl]methacrylamide,ethylene glycol dimethacrylate, PVP, alginate, and chitosan; a barrierlayer made of one or more hydrophobic components selected from the groupconsisting of: TRIS, SIGMA, fluorosiloxane macromer,mono-(3-methacryloxy-2-hydroxypropyloxy)propyl terminated, mono-butylterminated polydimethylsiloxane, mono-methacryloxypropyl terminatedpolydimethylsiloxane, and PMMA, configured to block the drug diffusionpaths from the reservoir domain to the ocular surface in the eye of thesubject; wherein the drug partitions between the reservoir domain andthe barrier layer, and the equilibrium drug solubility in the reservoirdomain is at least five folds higher than that in the barrier layer, andwhen the device is placed directly onto the ocular surface or in contactwith a tear film above the ocular surface in the eye, the deviceprovides controlled and sustained release of the drug at a constantrate.
 13. The transparent ocular device of claim 12, wherein the one ormore hydrophobic components are selected from the group consisting of:TRIS, SIGMA, fluorosiloxane macromer, and PMMA.
 14. The transparentocular device of claim 12, wherein the one or more hydrophiliccomponents are selected from the group consisting of: HEMA, DMA,ethylene glycol dimethacrylate, PVP, alginate, and chitosan.
 15. Thetransparent ocular device according to claim 12, wherein the one or morereservoir domains are in a form of a layer.
 16. The transparent oculardevice according to claim 12, wherein the one or more reservoir domainsare embedded within the barrier layer.
 17. The transparent ocular deviceaccording to claim 12, which is a contact lens, a scleral lens, anorthokeratology lens, or a corneal bandage.
 18. The transparent oculardevice of claim 2, wherein the one or more hydrophilic components areselected from the group consisting of: HEMA, DMA, ethylene glycoldimethacrylate, PVP, alginate, and chitosan.
 19. The transparent oculardevice of claim 2, wherein and the one or more hydrophilic componentsare selected from the group consisting of: DMA, ethylene glycoldimethacrylate, and PVP.